The need and demand for an accurate, non-invasive method for determining blood glucose level in patients is well documented. Barnes et al. (U.S. Pat. No. 5,379,764) disclose the necessity for diabetics to frequently monitor glucose levels in their blood. It is further recognized that the more frequent the analysis, the less likely there will be large swings in glucose levels. These large swings are associated with the symptoms and complications of the disease, whose long-term effects can include heart disease, arteriosclerosis, blindness, stroke, hypertension, kidney failure, and premature death. As described below, several systems have been proposed for the non-invasive measurement of glucose in blood. However, despite these efforts a lancet cut into the finger is still necessary for all presently commercially available forms of home glucose monitoring. This is believed so compromising to the diabetic patient that the most effective use of any form of diabetic management is rarely achieved.
The various proposed non-invasive methods for determining blood glucose level, discussed individually below, generally utilize quantitative infrared spectroscopy as a theoretical basis for analysis. Infrared spectroscopy measures the electromagnetic radiation (0.7-25 .mu.m) a substance absorbs at various wavelengths. Molecules do not maintain fixed positions with respect to each other, but vibrate back and forth about an average distance. Absorption of light at the appropriate energy causes the molecules to become excited to a higher vibration level. The excitation of the molecules to an excited state occurs only at certain discrete energy levels, which are characteristic for that particular molecule. The most primary vibrational states occur in the mid-infrared frequency region (i.e., 2.5-25 .mu.m). However, non-invasive analyte determination in blood in this region is problematic, if not impossible, due to the absorption of the light by water. The problem is overcome through the use of shorter wavelengths of light which are not as attenuated by water. Overtones of the primary vibrational states exist at shorter wavelengths and enable quantitative determinations at these wavelengths.
It is known that glucose absorbs at multiple frequencies in both the mid- and near-infrared range. There are, however, other infrared active analytes in the blood which also absorb at similar frequencies. Due to the overlapping nature of these absorption bands, no single or specific frequency can be used for reliable non-invasive glucose measurement. Analysis of spectral data for glucose measurement thus requires evaluation of many spectral intensities over a wide spectral range to achieve the sensitivity, precision, accuracy, and reliability necessary for quantitative determination. In addition to overlapping absorption bands, measurement of glucose is further complicated by the fact that glucose is a minor component by weight in blood, and that the resulting spectral data may exhibit a non-linear response due to both the properties of the substance being examined and/or inherent non-linearities in optical instrumentation.
Another problem encountered in non-invasive skin based measurements of standard medical blood analytes in order to replace the need to draw blood from the patient has been the inherent differences between the concentration of a given analyte in the blood and the same analyte in the overall skin tissue water. Much of the work toward a replacement for blood drawing has been focused on the measurement of blood glucose in diabetic patients who must lance themselves four to five times per day in order to measure their capillary blood glucose concentration and adjust insulin therapy and meals. In the case of the infrared measurement, the beam "interrogates" a tissue volume that is largely water (70-80%).
However, blood, which is also approximately 80% water, makes up less than 10% of the tissue volume. Since glucose is not made, but only disposed of, in skin, all of the glucose in the water that bathes cells (interstitial fluid) and that is inside cells comes from the blood vessels. That is, blood glucose must move out of the blood vessels and into the surrounding interstitial water and then into cellular elements. This effect is, of course, time dependent as well as dependent upon the gradients, relative juxtaposition of the compartments, as well as the relative blood flow to the tissue. In short, the relationship between blood and tissue glucose concentration is very complex and variable even in a single subject. Thus, an integrated or summed measurement of total tissue water glucose concentration is often very different from the concentration of glucose in the small blood vessels that make up a fraction of the total tissue volume.
Glucose concentration measurement of interstitial fluid (the usually clear fluid that bathes all cells outside of blood vessels) as a surrogate for direct blood glucose concentration is problematic for some of the same reasons. Instead of measuring all compartments as with spectroscopic techniques, only one compartment is measured. Again, since glucose is only degraded in the skin (not manufactured), the interstitial space must be "filled" with glucose by the local blood vessels. This is analogous to a dye being slowly dripped into a glass of water, the faster the dye is dripped, the faster it reaches a high concentration or dark color throughout the total volume. As with any filling process, this is time dependent. Time lags between the concentration of glucose in interstitial fluid and blood have been documented ranging from zero to 60 minutes with an average lag of 20 minutes. Thus, the fact that the glucose must move between the tissue and blood causes errors in both interstitial space glucose and total tissue glucose concentration measurements.
When measurements of total tissue or interstitial glucose concentration and blood glucose concentration are made concurrently, the two are correlated, but the tissue glucose concentrations lag behind the blood levels. Blood or serum glucose concentrations must be delayed in order to overlay the interstitial or total glucose concentration. When blood glucose concentration is changing rapidly as might be expected in a diabetic after a meal high in simple carbohydrates (sugars) or after an insulin injection, the delay is more obvious and the difference between the blood and the other two measurements is most pronounced. The error between the blood measurement and the total or interstitial measurements is highest.
This presents obvious problems with respect to using the surrogate methods for monitoring and basing therapy in diabetic patients. Given the concentration difference, determining whether a given technique is working based on infrequent, discrete measurements is nearly impossible. Without continuous measurements, it is difficult to determine whether the patient's blood glucose is in a steady state condition or is in a flux; increasing or decreasing.
The worst case scenario in diabetic glucose management would be a quickly falling blood glucose concentration. Such a situation could result following a large insulin injection, unopposed by either glucose production in the liver or carbohydrate uptake from food in the gut. If a tissue measurement were made it would inappropriately report a level which is higher than the actual blood glucose concentration. Thus, the patient would be unaware of their actual low blood glucose level. The result of very low blood glucose concentrations (below 40 mg/dl, 2.2 mmol) is often coma and even brain damage or death if the patient is not discovered in time for medical intervention. Thus, improving the agreement between blood and tissue measurements is desired.
A further common element to non-invasive glucose measuring techniques is the necessity for an optical interface between the body portion at the point of measurement and the sensor element of the analytical instrument. Generally, the sensor element must include an input element or means for irradiating the sample point with infrared energy. The sensor element must further include an output element or means for measuring transmitted or reflected energy at various wavelengths resulting from irradiation through the input element.
Robinson et al. (U.S. Pat. No. 4,975,581) disclose a method and apparatus for measuring a characteristic of unknown value in a biological sample using infrared spectroscopy in conjunction with a multivariate model that is empirically derived from a set of spectra of biological samples of known characteristic values. The above-mentioned characteristic is generally the concentration of an analyte, such as glucose, but also may be any chemical or physical property of the sample. The method of Robinson et al. involves a two-step process that includes both calibration and prediction steps. In the calibration step, the infrared light is coupled to calibration samples of known characteristic values so that there is differential attenuation of at least several wavelengths of the infrared radiation as a function of the various components and analytes comprising the sample with known characteristic value. The infrared light is coupled to the sample by passing the light through the sample or by reflecting the light from the sample. Absorption of the infrared light by the sample causes intensity variations of the light that are a function of the wavelength of the light. The resulting intensity variations at the at least several wavelengths are measured for the set of calibration samples of known characteristic values. Original or transformed intensity variations are then empirically related to the known characteristic of the calibration samples using a multivariate algorithm to obtain a multivariate calibration model. In the prediction step, the infrared light is coupled to a sample of unknown characteristic value, and the calibration model is applied to the original or transformed intensity variations of the appropriate wavelengths of light measured from this unknown sample. The result of the prediction step is the estimated value of the characteristic of the unknown sample. The disclosure of Robinson et al. is incorporated herein by reference.
Several of the embodiments disclosed by Robinson et al. are non-invasive and incorporate an optical interface having a sensor element. As depicted in FIGS. 5 and 6 of Robinson et al., the optical interface includes first, an input element and second, an output element. The input element is an infrared light source or near infrared light source. The input element interface with the sample or body portion containing blood to be tested includes transmitting the light energy or propagating the light energy to the surface of the skin via the air. The output element includes a detector which receives the transmitted or reflected light energy. The output interface with the sample also includes propagating the transmitted or reflected light through the air from the skin.
Wall et al. in PCT Application WO 92/17765 disclose a method for measuring glucose within a blood sample utilizing a radiation beam having a wavelength in the bandwidth of 1500 nm to 1700 nm, and a reference radiation source emitting a radiation beam having a wavelength in the bandwidth of 1200 to 1400 nm. Both beams pass through a test medium of blood to a detector arranged to detect and produce an output signal dependent upon the intensity of radiation beams impinging thereon. Wall et al. disclose that it is preferred that the blood sample be heated because it was found that if the temperature of the blood in the cuvette was elevated to around 40.degree. C., the amplitude of the light beam transmitted to a photodetector through the sample increased considerably. Wall et al. further state that for in vivo analysis, an electrically heated sleeve can be utilized as a finger-receiving cavity.
MacGregor et al. in PCT Application WO 93/07801 disclose a method and apparatus for determining non-invasively the presence and concentration of blood analytes such as glucose. The apparatus comprises a light source for producing a polychromatic light beam and means for modulating the polychromatic light beam, such that the modulation frequency is dependent upon the wavelength of light within the beam. The modulated light beam is caused to impinge upon a body part so that blood analytes interact with the light beam and perturb the spectral distribution of light within the beam. Spectral information is extracted from the resulting light beam by detecting the beam at a plurality of modulation frequencies. MacGregor et al. disclose that it is desirable to raise or lower the temperature of the body part to a constant temperature to minimize the variability in its spectral properties. It is disclosed that it is preferable to raise the body temperature, because the increasing temperature of the body part increases the amount of blood in the tissue and increases the strength of the pulsatile component of flow.
Robinson (U.S. Pat. No. 5,830,132) discloses a robust accurate non-invasive analyte monitor. The disclosure of Robinson is incorporated herein by reference. The method includes irradiating the tissue with infrared energy having at least several wavelengths in a given range of wavelengths so that there is differential absorption of at least some of the wavelengths by the tissue as a function of the wavelengths and the known characteristic, wherein the differential absorption causes intensity variations of the wavelengths incident from the tissue. The method further includes providing a first path through the tissue and a second path through the tissue, wherein the first path is optimized for a first sub-region of the range of wavelengths to maximize the differential absorption by at least some of the wavelengths in the first sub-region and then optimizing the second path for a second sub-region of the range to maximize the differential absorption by at least some of the wavelengths in the second sub-region. Robinson further discloses that the object of the invention is to measure blood analytes, therefore, maximizing the amount of blood in the tissue being irradiated is recognized as improving the measurement. The accuracy of non-invasive measurement is determined by its correlation to standard invasive blood measurements. To improve the stability and accuracy of the Robinson measurement, it is disclosed that a minimum sampling device should be thermostated so that the device does not act as a heat sink. It is further disclosed that the sampling device can be heated to an above normal tissue temperature to increase blood flow to the tissue area in contact with the device. The result is an increase in the vascular supply to the tissue and a corresponding increase in the blood content of the tissue. The end result of temperature regulation is taught as a reduction in spectral variation not associated with glucose and an improvement in measurement accuracy.
Barnes et al. (U.S. Pat. No. 5,379,764) disclose a spectrographic method for analyzing glucose concentration, wherein near infrared radiation is projected on a portion of the body, the radiation including a plurality of wavelengths, followed by sensing the resulting radiation emitted from the portion of the body as affected by the absorption of the body. The method disclosed includes pretreating the resulting data to minimize influences of offset and drift to obtain an expression of the magnitude of the sensed radiation as modified.
The sensor element disclosed by Barnes et al. includes a dual conductor fiber optic probe which is placed in contact or near contact with the skin of the body. The first conductor of the dual conductor fiber optic probe acts as an input element which transmits the near infrared radiation to the skin surface while in contact therewith. The second conductor fiber of the dual conductor probe acts as an output element which transmits the reflected energy or non-absorbed energy back to a spectrum analyzer. The optical interface between the sensor element and the skin is achieved by simply contacting the skin surface with the probe, and can include transmitting the light energy through air to the skin and through air back to the probe depending upon the degree of contact between the probe and skin. Irregularities in the skin surface and at the point of measurement will affect the degree of contact.
Dahne et al. (U.S. Pat. No. 4,655,225) disclose the employment of near infrared spectroscopy for non-invasively transmitting optical energy in the near infrared spectrum through a finger or earlobe of a subject. Also discussed is the use of near infrared energy diffusely reflected from deep within the tissues. Responses are derived at two different wavelengths to quantify glucose in the subject. One of the wavelengths is used to determine background absorption, while the other wavelength is used to determine glucose absorption.
The optical interface disclosed by Dahne et al. includes a sensor element having an input element which incorporates a directive light means which is transmitted through the air to the skin surface. The light energy which is transmitted or reflected from the body tissue as a measure of absorption is received by an output element. The interface for the output element includes transmitting the reflected or transmitted light energy through air to the detector elements.
Caro (U.S. Pat. No. 5,348,003) discloses the use of temporally-modulated electromagnetic energy at multiple wavelengths as the irradiating light energy. The derived wavelength dependence of the optical absorption per unit path length is compared with a calibration model to derive concentrations of an analyte in the medium.
The optical interface disclosed by Caro includes a sensor element having an input element, wherein the light energy is transmitted through a focusing means onto the skin surface. The focusing means may be near or in contact with the skin surface. The sensor element also includes an output element which includes optical collection means which may be in contact with the skin surface or near the skin surface to receive light energy which is transmitted through the tissue. Again, a portion of the light energy is propagated through air to the skin surface and back to the output element due to non-contact with the sensor and irregularities in the skin surface.
Problems with the optical interface between the tissue and the instrument have been recognized. In particular, optical interface problems associated with coupling light into and back out of the tissue were recognized by Ralf Marbach as published in a thesis entitled "MeBverfahren zur IR-spektroskopishen Blutglucose Bestimmung" (English translation "Measurement Techniques for IR Spectroscopic Blood Glucose Determination"), published in 1993.
Marbach states that the requirements of the optical accessory for measurement of the diffuse reflection of the lip are:
1) High optical "throughput" for the purpose of optimizing the S/N ratio of the spectra, and
2) Suppression of the insensitivity to Fresnel or specular reflection on the skin surface area.
The measurement accessory proposed by Marbach attempts to meet both requirements through the use of a hemispherical immersion lens. The lens is made out of a material which closely matches the refractive index of tissue, calcium fluoride. As stated by Marbach, the important advantages of the immersion lens for transcutaneous diffuse reflection measurements are the nearly complete matching of the refraction indices of CaF.sub.2 and skin and the successful suppression of the Fresnel reflection.
Calcium fluoride, however is not an ideal index match to tissue, having an index of 1.42, relative to that of tissue, at approximately 1.38. Thus, an index mismatch occurs at the lens to tissue interface assuming complete contact between the lens and tissue. The optical efficiency of the sampling accessory is further compromised by the fact that the lens and the tissue will not make perfect optical contact due to roughness of the tissue. The result is a significant refractive index mismatch where the light is forced to travel from the lens (N=1.42) to air (N=1.0) to tissue (N=1.38). Thus, the inherent roughness of tissue results in small air gaps between the lens and the tissue, which decrease the optical throughput of the system, and subsequently compromise the performance of the measurement accessory.
The magnitude of the problem associated with refractive index mismatch is a complicated question. First, a fraction of light, which would otherwise be available for spectroscopic analysis of blood analytes, gets reflected at the mismatch boundary and returns to the input or collection optical system without interrogating the sample. The effect is governed by the Fresnel Equation: ##EQU1##
For normally incident, randomly polarized light, where N and N' are the refractive indices of the two media. Solving for the air/CaF.sub.2 interface gives an R=0.03, or a 3% reflection. This interface must be traversed twice, leading to a 6% reflected component which does not interrogate the sample. These interface mismatches are multiplicative. The fraction of light successfully entering the tissue then must be considered. In some regions of the spectrum, for instance, under a strong water band, almost all of the transmitted light gets absorbed by the tissue. The result is that this seemingly small reflected light component from the refractive index mismatch can virtually overwhelm and obscure the desired signal from the sample.
Finally, it is useful to consider the critical angle effect as light attempts to exit the tissue. Tissue is highly scattering and so a light ray which launches into tissue at normal incidence may exit the tissue at a high angle of incidence. If the coupling lens is not in intimate contact with the tissue, these high angle rays will be lost to total internal reflection. The equation which defines the critical angle, or the point of total internal reflection, is as follows: ##EQU2##
When light is propagating through a higher index material like tissue (N'=1.38) and approaching an interface with lower refractive index like air (N=1.0), a critical angle of total internal reflection occurs. Light approaching such an interface at greater than the critical angle will not propagate into the rarer medium (air), but will totally internally reflect back into the tissue. For the aforementioned tissue/air interface, the critical angle is 46.4. No light steeper than this angle would escape. Intimate, optical contact is therefore essential to efficient light capture from tissue.
As detailed above, each of the prior art apparatus for non-invasively measuring glucose concentration utilize a sensor element. Each sensor element includes an input element and an output element. The optical interface between the input element, output element and the skin surface of the tissue to be analyzed in each apparatus is similar. In each instance, the input light energy is transmitted through air to the surface or potentially through air due to a gap in the contact surface between the input sensor and the skin surface. Likewise, the output sensor receives transmitted or reflected light energy via transmission through air to the output sensor, or potentially through a gap between the sensor element and the skin surface even though attempts are made to place the output sensor in contact with the skin. It is believed that the optical interfaces disclosed in the prior art affect the accuracy and consistency of the data acquired utilizing the prior art methods and apparatus. Thus, the accuracy of these methods for non-invasively measuring glucose are compromised.
Wu et al. (U.S. Pat. No. 5,452,723) disclose a method of spectrographic analysis of a tissue sample, which includes measuring the diffuse reflectance spectrum, as well as a second selected spectrum, such as fluorescence, and adjusting the spectrum with the reflectance spectrum. Wu et al. assert that this procedure reduces the sample-to-sample variability. Wu et al. disclose the use of an optical fiber as an input device that is bent at an acute angle so that incident light from the fiber impinges on an optically smooth surface of an optical coupling medium. The optical coupling medium is indexed matched to the tissue so that little or no specular reflection occurs at the interface between the catheter and the tissue. Wu et al. further disclose that the catheter can be used in contact or non-contact modes with the tissue. In contact mode, the end of the catheter is placed in direct contact with the tissue to accomplish index matched optical coupling. Thus, the optical coupling medium of Wu et al. is a solid end portion on the optical fiber. Wu et al. further disclose that the catheter can be used in a non-contact mode, wherein the gap left between the end of the catheter and the tissue can be filled with an index-matched fluid to prevent specular reflections. The only criteria disclosed throughout the Wu et al. specification for the fluid is that it is index matched to prevent specular reflections, which is only one aspect of an optimum optical interface for spectrographic analysis of an analyte in blood.
Accordingly, the need exists for a method and apparatus for non-invasively measuring glucose and other analyte concentrations in blood which accounts for or corrects problems associated with differences in analyte concentration in the various fluid compartments that comprise a tissue area or volume being tested. Further, there is a need for an apparatus and method to determine whether analyte concentrations are rising, falling or at equilibrium along with an indication of the rate of change in order to optimize treatment in response to the data. A preferred apparatus should incorporate an improved optical interface. The optical interface should produce consistent repeatable results so that the analyte concentration can be accurately calculated from a model such as that disclosed by Robinson et al. The optical interface should maximize both the input and output light energy from the source into the tissue and from the tissue back to the output sensor. The detrimental effects of gaps due to irregularities in the surface of the skin or the presence of other contaminants should be reduced or eliminated. Means should also be provided to guarantee that such optimized interface is achieved each time a user is coupled to the device for analysis.
The present invention addresses these needs as well as other problems associated with existing methods for non-invasively measuring glucose concentration in blood utilizing infrared spectroscopy and the optical interface associated therewith. The present invention also offers further advantages over the prior art and solves problems associated therewith.